Medical device for delivering a therapeutic substance and method therefor

ABSTRACT

A device useful for localized delivery of a therapeutic material is provided. The device includes a structure including a porous material; and a water-insoluble salt of a therapeutic material dispersed in the porous material. The water-insoluble salt is formed by contacting an aqueous solution of a therapeutic salt with a heavy metal water-soluble salt dispersed throughout a substantial portion of the porous material. The porous material can be made of a polymer other than fibrin with fibrin incorporated into the pores, which can be the only layer of polymeric material on the medical device (e.g., stent). A new method for preparing a porous polymer material on a medical device.

BACKGROUND OF THE INVENTION

[0001] This invention relates to a medical device employing atherapeutic substance as a component thereof. For example, in anarterial site treated with percutaneous transluminal coronaryangioplasty therapy for obstructive coronary artery disease atherapeutic antithrombogenic substance such as heparin may be includedwith a device and delivered locally in the coronary artery. Alsoprovided is a method for making a medical device capable of localizedapplication of therapeutic substances. This invention also relates to amedical device, particularly a stent, having a porous polymeric filmwith fibrin incorporated therein for enhanced biocompatibility, with orwithout a therapeutic substance as a component thereof.

[0002] Medical devices which serve as substitute blood vessels,synthetic and intraocular lenses, electrodes, catheters and the like inand on the body or as extracorporeal devices intended to be connected tothe body to assist in surgery or dialysis are well known. For example,intravascular procedures can bring medical devices into contact with thepatient's vasculature. In treating a narrowing or constriction of a ductor canal percutaneous transluminal coronary angioplasty (PTCA) is oftenused with the insertion and inflation of a balloon catheter into astenotic vessel. Other intravascular invasive therapies includeatherectomy (mechanical systems to remove plaque residing inside anartery), laser ablative therapy and the like. However, this use ofmechanical repairs can have adverse consequences for the patient. Forexample, restenosis at the site of a prior invasive coronary arterydisease therapy can occur. Restenosis, defined angiographically, is therecurrence of a 50% or greater narrowing of a luminal diameter at thesite of a prior coronary artery disease therapy, such as a balloondilatation in the case of PTCA therapy. In particular, an intra-luminalcomponent of restenosis develops near the end of the healing processinitiated by vascular injury, which then contributes to the narrowing ofthe luminal diameter. This phenomenon is sometimes referred to as“intimal hyperplasia.” It is believed that a variety of biologic factorsare involved in restenosis, such as the extent of the injury, platelets,inflammatory cells, growth factors, cytokines, endothelial cells, smoothmuscle cells, and extracellular matrix production, to name a few.

[0003] Attempts to inhibit or diminish restenosis often includeadditional interventions such as the use of intravascular stents and theintravascular administration of pharmacological therapeutic agents.Examples of stents which have been successfully applied over a PTCAballoon and radially expanded at the same time as the balloon expansionof an affected artery include the stents disclosed in U.S. Pat. No.4,733,665 issued to Palmaz, U.S. Pat. No. 4,800,882 issued to Gianturcoand U.S. Pat. No. 4,886,062 issued to Wiktor. Also, such stentsemploying therapeutic substances such as glucocorticoids (e.g.dexamethasone, beclamethasone), heparin, hirudin, tocopherol,angiopeptin, aspirin, ACE inhibitors, growth factors, oligonucleotides,and, more generally, antiplatelet agents, anticoagulant agents,antimitotic agents, antioxidants, antimetabolite agents, andanti-inflammatory agents have been considered for their potential tosolve the problem of restenosis. Such substances have been incorporatedinto a solid composite with a polymer in an adherent layer on a stentbody with fibrin in a separate adherent layer on the composite to form atwo layer system. The fibrin is optionally incorporated into a porouspolymer layer in this two layer system.

[0004] Another concern with intravascular and extracorporeal proceduresis the contact of biomaterials with blood which can trigger the body'shemostatic process. The hemostatic process is normally initiated as thebody's response to injury. When a vessel wall is injured, plateletsadhere to damaged endothelium or exposed subendothelium. Followingadhesion of the platelets, these cells cohere to each other preparatoryto aggregation and secretion of their intracellular contents.Simultaneously there is activation, probably by electrostatic charge ofthe contact factors, of the coagulation cascade. The sequentialstep-wise interaction of these procoagulant proteins results in thetransformation of soluble glycoproteins into insoluble polymers, whichafter transamidation results in the irreversible solid thrombus.

[0005] Immobilization of polysaccharides such as heparin to biomaterialshas been used to improve bio- and hemocompatibility of implantable andextracorporeal devices. The mechanism responsible for reducedthrombogenicity of heparinized materials is believed to reside in theability of heparn to speed up the inactivation of serine proteases(blood coagulation enzymes) by AT-III. In the process, AT-III forms acomplex with a well defined pentasaccharide sequence in heparin,undergoing a conformational change and thus enhancing the ability ofAT-III to form a covalent bond with the active sites of serine proteasessuch as thrombin. The formed TAT-complex then releases from thepolysaccharide, leaving the heparin molecule behind for a second roundof inactivation.

[0006] Usually, immobilization of heparin to a biomaterial surfaceconsists of activating the material in such a way that coupling betweenthe biomaterial and functional groups on the heparin (—COOH, —OH, —NH₂)can be achieved. For example, Larm presented (in U.S. Pat. No.4,613,665) a method to activate heparin via a controlled nitrous aciddegradation step, resulting in degraded heparin molecules of which apart contains a free terminal aldehyde group. Heparin in this form canbe covalently bound to an aminated surface in a reductive aminationprocess. Although the molecule is degraded and as a result shows lesscatalytic activity in solution, the end point attachment of this type ofheparin to a surface results in true anti-thrombogenicity due to theproper presentation of the biomolecule to the surface. In this fashion,the molecule is freely interacting with AT-III and the coagulationenzymes, preventing the generation of thrombi and microemboli.

[0007] However, the attachment and delivery of therapeutic substancessuch as heparin can involve complicated and expensive chemistry. It istherefore an object of the present invention to provide a medical devicehaving a biocompatible, blood-contacting surface with an activetherapeutic substance at the surface and a simple, inexpensive methodfor producing such a surface. It is also an object of the presentinvention to provide a medical device having a porous material withfibrin incorporated therein, optionally with an active therepeuticsubstance at the blood-contacting surface. It is also a further objectof the present invention to provide a medical device, such as anintravascular stent, having a porous polymeric film adhered to themedical device body with fibrin incorporated therein for enhancedbiocompatibility.

SUMMARY OF THE INVENTION

[0008] This invention relates to a medical device having ablood-contacting surface with a therapeutic substance thereon.Preferably, the device according to the invention is capable of applyinga highly localized therapeutic material into a body lumen to treat orprevent injury. The term “injury” means a trauma, that may be incidentalto surgery or other treatment methods including deployment of a stent,or a biologic disease, such as an immune response or cell proliferationcaused by the administration of growth factors. In addition, the methodsof the invention may be performed in anticipation of “injury” as aprophylactic. A prophylactic treatment is one that is provided inadvance of any symptom of injury in order to prevent injury, preventprogression of injury or attenuate any subsequent onset of a symptom ofsuch injury.

[0009] In accordance with the invention, a device for delivery oflocalized therapeutic material includes a structure including a porousmaterial and a plurality of discrete particles of a water-insoluble saltof the therapeutic material dispersed throughout a substantial portionof the porous material. Preferably, the device is capable of beingimplanted in a body so that the localized therapeutic agent can bedelivered in vivo, typically at a site of vascular injury or trauma.More preferably, the porous material is also biocompatible, sufficientlytear-resistant and nonthrombogenic.

[0010] The porous material may be a film on at least a portion of thestructure or the porous material may be an integral portion of thestructure. Preferably, the porous material is selected from the group ofa natural hydrogel, a synthetic hydrogel, TEFLON(polytetrafluoroethylene), silicone, polyurethane, polysulfone,cellulose, polyethylene, polypropylene, polyamide, polyester, and acombination of two or more of these materials. Examples of naturalhydrogels include fibrin, collagen, elastin, and the like.

[0011] Alternatively, the porous material may have fibrin incorporatedtherein. Although this material preferably has a therapeutic agent alsoincorporated therein, this is not necessary for enhancedbiocompatibility. Thus, in one embodiment, the present inventionprovides a medical device, preferably, an intravascular stent, thatincludes a porous polymer film with fibrin incorporated within thepores, optionally with a therapeutic substance also incorporated withinthe pores.

[0012] The therapeutic agent preferably includes an antithromboticmaterial. More preferably, the antithrombotic material is a heparin orheparin derivative or analog. Also preferably, the insoluble salt of thetherapeutic material is one of the silver, barium or calcium salts ofthe material.

[0013] The structure of the device can be adapted for its intendedextracorporeal or intravascular purpose in an internal human body site,such as an artery, vein, urethra, other body lumens, cavities, and thelike or in an extracorporeal blood pump, blood filter, blood oxygenatoror tubing. In one aspect of the invention, the shape is preferablygenerally cylindrical, and more preferably, the shape is that of acatheter, a stent, or a guide wire.

[0014] In another aspect of the invention, an implantable device capableof delivery of a therapeutic material includes a structure comprising aporous material; and a plurality of discrete particles comprising aheavy metal water-soluble salt dispersed throughout a substantialportion of the porous material. Preferably, the heavy metalwater-soluble salt is selected from the group of AgNO₃, Ba(NO₃)₂, BaCl₂,and CaCl₂. The amount of water-soluble salt dispersed throughout aportion of the porous material determines the total amount oftherapeutic material that can be delivered once the device is implanted.

[0015] The invention provides methods for manufacturing medical devices.Specifically, the invention provides a method for coating a medicaldevice with a porous polymer (film or coating). The method includes:placing the medical device in a mold; placing a solution of a polymer inthe mold with the medical device; wherein the solution of the polymerincludes a solvent capable of phase separating from the polymer at atemperature below the freezing point of the solvent; cooling thesolution of the polymer in the mold to a temperature below the freezingpoint of the solvent until a first fraction of particulate material isformed by solidification and phase separation of the solvent from thepolymer and is dispersed within solidified polymer; cooling the solutionfurther and at a faster rate than in the first cooling step to form asecond fraction of particulate material dispersed within the solidifiedpolymer, wherein the second fraction of particulate material has asmaller particle size than the first fraction; and removing theparticulate material from the polymer to form pores therein. Preferably,the medical device is a stent and the solution includes a polyurethanedissolved in dioxane.

[0016] The invention also provides methods for making an implantabledevice which includes therapeutic materials. In one embodiment, a methodof the invention includes loading a structure comprising a porousmaterial with a heavy metal water-soluble salt dispersed throughout asubstantial portion of the porous material, sterilizing the loadedstructure, and packaging for storage and, optionally, delivery of thesterilized loaded structure. Preferably, the method of the inventionfurther includes substantially contemporaneously loading of a watersoluble therapeutic material, wherein a water insoluble salt of thetherapeutic material is produced throughout a substantial portion of theporous material of the structure. “Substantially contemporaneously,”means that the step of loading a water soluble therapeutic materialoccurs at or near a step of positioning the device proximate to adesired area, i.e., at or near the surgical arena prior toadministration to or implantation in, a patient. More preferably, thewater insoluble salt of the therapeutic material is dispersed throughouta substantial portion of the porous material.

[0017] In another aspect of the invention, a method includes loading astructure comprising a porous material with a heavy metal water-solublesalt dispersed throughout a substantial portion of the porous material;loading a water soluble therapeutic material, wherein a water insolublesalt of the therapeutic material is produced in a substantial portion ofthe porous material of the structure; and packaging for delivery of theloaded structure.

[0018] Thus, the methods for making an implantable device to deliver atherapeutic material and device in vivo, or in an extracorporeal circuitin accordance with the invention, are versatile. A therapeutic materialmay be loaded onto a structure including a porous material at any numberof points between, and including, the point of manufacture and the pointof use. As a result of one method, the device can be stored andtransported prior to incorporation of the therapeutic material. Thus,the end user can select the therapeutic material to be used from a widerrange of therapeutic agents.

BRIEF DESCRIPTION OF THE DRAWINGS

[0019]FIG. 1 is an elevational view of one embodiment of a deviceaccording to the invention with a balloon catheter as a mode of deliveryof the device;

[0020]FIG. 2 is an elevational view of another embodiment of a deviceaccording to the invention with a balloon catheter as a mode of deliveryof the device;

[0021]FIG. 3 is a flow diagram schematically illustrating methodsaccording to the invention;

[0022]FIG. 4 is a photograph taken from a scanning electron microscopeof a surface showing the insoluble therapeutic material according to theinvention;

[0023]FIGS. 5a and 5 b are photographs showing the histologicalcomparison between a stent heparinized according to the presentinvention (5 a) and a control stent (5 b); and

[0024]FIGS. 6 and 7 are photographs showing different magnifications ofan artery wall with an expanded stent therein having a porous polymerfilm with fibrin incorporated therein.

DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS

[0025] One of the more preferred configurations for a device accordingto the invention is a stent for use in artery/vascular therapies. Theterm “stent” refers to any device capable of being delivered by acatheter and which, when placed into contact with a portion of a wall ofa lumen to be treated, will also deliver localized therapeutic materialat a luminal or blood-contacting portion of the device. A stenttypically includes a lumen wall-contacting surface and a lumen-exposedsurface. Where the stent is shaped generally cylindrical or tube-like,including a discontinuous tube or ring-like structure, the lumen-wallcontacting surface is the surface in close proximity to the lumen wallwhereas the lumen-exposed surface is the inner surface of thecylindrical stent. The stent can include polymeric or metallic elements,or combinations thereof, onto which a porous material is applied. Forexample, a deformable metal wire stent is useful as a stent framework ofthis invention, such as that described in U.S. Pat. No. 4,886,062 toWiktor, which discloses preferred methods for making a wire stent. Othermetallic stents useful in this invention include those of U.S. Pat. No.4,733,655 to Palmaz and U.S. Pat. No. 4,800,882 to Gianturco.

[0026] Referring now to FIG. 1, the stent 20 comprises a stent framework22 and a porous material coating 24. The stent framework 22 isdeformable and can be formed from a polymeric material, a metal or acombination thereof. A balloon 15 is positioned in FIG. 1 adjacent thelumen-exposed surface of the stent to facilitate delivery of the stent.The stent 20 can be modified to increase or to decrease the number ofwires provided per centimeter in the stent framework 22. Similarly, thenumber of wire turns per centimeter can also be modified to produce astiffer or a more flexible stent framework.

[0027] Polymeric stents can also be used in this invention. The polymerscan be nonbioabsorbable or bioabsorbable in part, or total. Stents ofthis invention can be completely nonbioabsorbable, totally bioabsorbableor a composite of bioabsorbable polymer and nonabsorbable metal orpolymer. For example, another stent suitable for this invention includesthe self-expanding stent of resilient polymeric material as disclosed inInternational Publication No. WO 91/12779.

[0028] Nonbioabsorbable polymers can be used as alternatives to metallicstents. The stents of this invention should not substantially induceinflammatory and neointimal responses. Examples of biostablenonabsorbable polymers that have been used for stent construction withor without metallic elements include polyethylene terephthalate (PET),polyurethane urea and silicone (for example, see van Beusekom et al.,Circulation, 86(supp. I):I-731, 1992 and Lincoff et al., J. Am. CollCardiol., 21(supp. 1):335A, 1994. Although the porous material is shownas a coating 24, it is to be understood that, for the purposes of thisinvention, the porous material can be incorporated into the material ofthe stent.

[0029] Referring to FIG. 2, an alternative stent 30 is shown. The stentframework 34 is affixed with a film of a porous material 32. This can beaccomplished by wrapping the film 32 around the stent framework 34 andsecuring the film 32 to the framework 34 (i.e., the film is usuallysufficiently tacky to adhere itself to the framework but a medical gradeadhesive could also be used if needed) so that the film 32 will stay onthe balloon 36 and framework 34 until it is delivered to the site oftreatment. The film 32 is preferably wrapped over the framework withfolds or wrinkles that will allow the stent 30 to be readily expandedinto contact with the wall of the lumen to be treated. Alternatively,the film 32 can be molded to the stent framework 34 such that theframework 34 is embedded within the film 32. Preferably, the film 32 islocated on a lumen-wall contacting surface 33 of the stent framework 34such that therapeutic material is substantially locally delivered to alumen wall, for example, an arterial wall membrane (not shown).

[0030] Porous Material

[0031] As mentioned above, the device according to the invention isgenerally a structure including a porous material. In one embodiment,the porous material is a film on at least a portion of the structure. Inanother embodiment, the porous material is an integral portion of thestructure. Preferably, the porous material is biocompatible, andsufficiently tear-resistant and nonthrombogenic. More preferably, theporous material is selected from the group of a natural hydrogel, asynthetic hydrogel, TEFLON (polytetrafluoroethylene), silicone,polyurethane, polysulfone, cellulose, polyethylene, polypropylene,polyamide, polyester, and a combination of two or more of thesematerials. Examples of natural hydrogels include fibrin, collagen,elastin, and the like. In materials which do not include pores in theirusual structural configurations, pores between one micrometer indiameter or as large as 1000 micrometers in diameter can be introducedby conventional means such as by introducing a solvent solubleparticulate material into the desired structure and dissolving theparticulate material with a solvent. However, no particular pore size iscritical to this invention.

[0032] In a preferred embodiment, the porous material is made of apolymer other than fibrin (i.e., a non-fibrin porous material) and hasfibrin incorporated within the pores. Typically, and preferably, theporous material is in the form of a sheet material of a syntheticbiostable polymer. Such synthetic biostable polymers include silicone,polyurethane, polysulfone, cellulose, polyethylene, polypropylene,polyamide, polyester, polytetrafluoroethylene, and combinations thereof.

[0033] If the porous material is in the form of a porous sheet or film,it can be made by a variety of methods. These methods can include, forexample, forming the films using a solid particulate material that canbe substantially removed after the film is formed, thereby formingpores. By using a solid particulate material during film formation thesize of the pores can, to some extent, be controlled by the size of thesolid particulate material being used. The particulate material canrange from less than about 1 micrometer in diameter to about 1000micrometers, preferably about 1 micrometer to about 100 micrometers,more preferably about 25 micrometers to about 60 micrometers. Foruniformity of pores, the particulate material can be screened throughsuccessively finer mesh sieves, e.g., through 100, 170, 270, 325, 400,and 500 mesh analytical grade stainless steel mesh sieves, to produce adesired range of particle sizes.

[0034] In one method according to the present invention, a porouspolyurethane sheet material (i.e., film) can be made by dissolving apolyether urethane in an organic solvent such as 1-methyl-2-pyrrolidone;mixing into the resulting polyurethane solution a crystalline,particulate material like a salt or sugar that is not soluble in thesolvent; casting the solution with particulate material into a thinfilm; and then applying a second solvent, such as water, to dissolve andremove the particulate material, thereby leaving a porous sheet. Such amethod is disclosed, for example, in U.S. Pat. Nos. 5,599,352 and5,591,227, both issued to Dinh et al. A portion of the particulatematerial may remain within the film. As a result, it is preferred thatthe solid particulate material be biocompatible.

[0035] Although films for stent bodies according to the presentinvention can be manufactured separately from the support structure ofthe stent and attached to the support structure after formation,preferred methods include forming the films directly on the supportstructure such that the support structure is at least partially,preferably completely, encapsulated by the film. In one such methoddisclosed in International Publication No. WO 97/07973, a stent isplaced on a mandrel. A particulate material is then applied to themandrel and stent such that it is lightly adhered to the mandrel. Theparticulate material should be readily soluble in a solvent which willnot also dissolve the polymer chosen for the film. For example,crystalline sodium bicarbonate is a water soluble material that can beused as the particulate material. A nonaqueous liquid, preferably asolvent for the polymer film material, can be applied to the mandrelbefore applying the particulate material in order to retain more of theparticulate material on the mandrel. For example, when a polyurethane isto be used for the film material, the solvent 1-methyl-2-pyrrolidinone(NMP) can be used to wet the surface of the mandrel before theapplication of particulate material. Preferably, the mandrel iscompletely dusted with the particulate in the portions of the mandrel tobe coated with the polymer film. This can be accomplished by dipping themandrel in NMP, allowing it to drain vertically for a few seconds andthen dusting the sodium bicarbonate onto the mandrel while rotating ithorizontally until no further bicarbonate particles adhere. Excessparticulate material can be removed by gently tapping the mandrel.

[0036] Coating with polymer may proceed immediately followingapplication of the particulate material. A polymer is provided in adilute solution and is applied to the particle-coated stent and mandrel.For example, polyurethane can be dissolved in NMP to make a 10%solution. Gel particles and particulate impurities can be removed fromthe solution by use of a clinical centrifuge. The polymer solution canbe applied by dipping the mandrel into the solution and letting thesolvent evaporate. With the solution of polyurethane and NMP, a singledip in the solution can provide a film of adequate thickness. To assistin the formation of communicating passageways through the polymerbetween the blood-contacting surface and the lumen-contacting surface,additional sodium bicarbonate particles are preferably dusted onto thepolymer solution immediately after the dipping operation and before thepolymer solution has dried. Excess particulate material can be removedby gently tapping the mandrel. To precipitate and consolidate thepolyurethane film on the stent, it can be dipped briefly (about 5minutes) in water and then rolled gently against a wetted surface, suchas a wet paper towel. The stent assembly can then be placed into one ormore water baths over an extended period (e.g., 8 hours) to dissolve andremove the sodium bicarbonate. After drying in air at temperatures fromabout 20° C. to about 50° C., the film then can be trimmed to match thecontour of the wire.

[0037] In yet another method, a solvent in which the polymer (i.e., afilm-forming polymer) is soluble that is capable of phase separatingfrom the polymer at a reduced temperature can be used to prepare aporous polymer film. In this method, the stent or other medical deviceis placed in a cavity of a mold designed for forming a film around thestent, similar to that disclosed in U.S. Pat. No. 5,510,077 to Dinh etal. A solution of the desired polymer, such as polyurethane, dissolvedin a solvent, such as dioxane, is added to the mold. The temperature ofthe solution is then reduced to a temperature at which the solventfreezes and phase separates from the polymer as it forms a film, therebyforming particulate material (i.e., frozen solvent particles) in situ.Typically, for polyurethane in dioxane, this is a temperature of about−70° C. to about 3° C. The composition is then immersed in an ice coldwater bath (at about 3° C.) for a few days to allow the dioxane todissolve into the ice cold water, thereby forming pores. The number andsize of the pores can be controlled by the concentration of the polymerand the freezing temperature. A method similar to this is disclosed inLiu et al., J. Biomed. Mater. Res., 26, 1489 (1992). This method can beimproved on by using a two-step freezing process described herein. In afirst step, the mixture is cooled slowly to create a first fraction ofparticulate material (i.e., frozen solvent particles) dispersed withinsolidified polymer. In a second step, the mixture is cooled further (andmore quickly) to create a second fraction of particulate material ofsmaller size dispersed within solidified polymer. In this way, a widerrange of pore sizes can be formed with greater control. This modifiedmethod is further described in Example 6.

[0038] Therapeutic Material

[0039] The therapeutic material used in the present invention could bevirtually any therapeutic substance which possesses desirabletherapeutic characteristics and which can be provided in both watersoluble and water insoluble salts and which have bioactivity as aninsoluble salt. For example, antithrombotics, antiplatelet agents,antimitotic agents, antioxidants, antimetabolite agents,anti-inflammatory agents and radioisotopes could be used. “Insolublesalt” or “water insoluble salt” of the therapeutic substance as setforth herein, means that the salt formed has a relatively poorsolubility in water such that it will not readily disperse from thepores of the device. In particular, anticoagulant agents such asheparin, heparin derivatives and heparin analogs could be used toprevent the formation of blood clots on the device. Also,water-insoluble radioactive salts such as Agl¹³⁵, BaS³⁵O₄, and(Ca)₃(P³²O₄)₂ could be used for application of radiotherapy to a bodylumen or blood.

[0040] Preferably, the water-insoluble salt of the therapeutic materialis formed by a heavy metal water-soluble salt interacting with anaqueous solution of the therapeutic material. In the present invention,the heavy metal water-soluble salt is dispersed throughout a substantialportion of the porous material. Preferably, the heavy metalwater-soluble salt is selected from the group of AgNO₃, Ba(NO₃)₂, BaCl₂,CaCl₂, and a mixture thereof. The amount of water-soluble salt dispersedthroughout a portion of the porous material determines the ultimateamount of therapeutic material capable of being administered once thedevice is implanted.

[0041] Fibrin and Methods of Incorporation into Porous Polymer

[0042] The term “fibrin” herein means naturally occurring polymer offibrinogen that arises during blood coagulation. It is an insoluble,crosslinked polymer generated by the action of thrombin on fibrinogen.Fibrinogen has three pairs of polypeptide chains (ALPHA 2-BETA 2-GAMMA2) covalently linked by disulfide bonds with a total molecular weight ofabout 340,000. Fibrinogen is converted to fibrin through proteolysis bya fibrinogen-coagulating protein, such as thrombin, reptilase, orancrod.

[0043] Methods of making fibrin and forming it into implantable devicesare well known in the art. See, for example, U.S. Pat. Nos. 4,548,736(Muller et al.) and 3,523,807 (Gerendas), and European PatentApplication 0 366 564. In one method, fibrin is formed by contactingfibrinogen with a fibrinogen-coagulating protein, such as thrombin,reptilase, or ancrod. Preferably, the fibrinogen andfibrinogen-coagulating protein (e.g., thrombin) used to make fibrin isfrom the same animal or human species as that in which the medicaldevice (e.g., stent) of the present invention will be implanted in orderto avoid cross-species immune reactions. The resulting fibrin can alsobe subjected to heat treatment at about 150° C. for 2 hours in order toreduce or eliminate antigenicity.

[0044] Preferably, the fibrinogen used to make the fibrin is abacteria-free and virus-free fibrinogen such as that described in U.S.Pat. No. 4,540,573 (Neurath et al.). The fibrinogen is preferably usedin solution at a concentration of at least about 10 mg/ml, morepreferably at least about 26 mg/ml, and no greater than about 50 mg/ml.The pH of the solution is preferably about 5.8 to about 9.0 with anionic strength of about 0.05 to about 0.45. The fibrinogen solution caninclude pure fibrinogen, although preferably the solution also includesproteins and enzymes such as albumin, fibronectin, Factor XIII,plasminogen, antiplasmin, Antithrombin HI, and the like. Mostpreferably, the fibrinogen is cryoprecipitated fibrinogen, which caninclude hundreds of proteins and enzymes, as disclosed in Spotmitz etal., The American Surgeon, 53, 460-462 (1987).

[0045] The fibrinogen solution also preferably includes afibrinogen-coagulating protein, such as thrombin. Alternatively,however, the fibrinogen-coagulating protein solution can be appliedafter the fibrinogen has been applied. This fibrinogen-coagulatingprotein solution may or may not include fibrinogen. The thrombin ispreferably used in solution at a concentration of at least about 1 NIHunit/ml, and no greater than about 120 NIH units/ml. Calcium ions mayalso be present in the thrombin solution to enhance mechanicalproperties and biostability of the device. If used, they are preferablypresent in a concentration of about 0.02 M to about 0.2 M.

[0046] Preferably, the coagulating effect of any residual coagulationprotein in the fibrin should be neutralized before employing it in amedical device, such as a stent in order to prevent clotting at thefibrin interface with blood after implantation. This can beaccomplished, for example, by treating the fibrin with irreversiblecoagulation inhibitor compounds or heat after polymerization. Forexample, hirudin or D-phenylalanyl-propylarginine chloromethyl ketone(PPACK) can be used for this purpose. Anticoagulants, such as heparin,can also be added to reduce the possibility of further coagulation.

[0047] The porous sheet can then be placed into a fibrinogen solution inorder to fill the pores with fibrinogen, followed by application of asolution of thrombin and fibrinogen to the surface of the sheet materialto establish a fibrin matrix that occupies both the surface of the sheetand the pores of the sheet. Alternatively, the thrombin can be includedwithin the first fibrinogen solution. If desired, ultrasonics, vacuum,and/or pressure can be used to ensure that the fibrinogen applied to thesheet is received into the pores.

[0048] Methods of Making an Implantable Device Having a TherapeuticSubstance

[0049] Referring now to FIG. 3, a structure having a porous material isloaded with a heavy metal water-soluble salt. Preferably, this stepincludes contacting, more preferably immersing, the structure with anaqueous solution of the heavy metal water-soluble salt, as describedabove. Preferably, the heavy metal water-soluble salt is dispersedthroughout a substantial portion of the porous material. This may beassisted by degassing the pores of the structure by such techniques asultrasound or vacuum degassing. The resulting stature can now besterilized, packaged and, optionally, stored until use.

[0050] In one embodiment of the invention, a sterilized structure isshipped or delivered to the relevant consumer. The structure issubstantially contemporaneously loaded with a water soluble therapeuticmaterial. Preferably, the loading of the therapeutic material includescontacting, more preferably immersing, the porous material in an aqueoussolution comprising a salt of the therapeutic material, as describedabove. Again, degassing of the device can help to bring the therapeuticmaterial into the pores. A water-insoluble therapeutic salt is therebyformed within the porous material. Examples of aqueous radioactive saltsolutions for radiotherapy include Nal¹²⁵, K₂S³⁵O₄, Na₂S³⁵O₄, andNa₃P³²O₄, to name a few.

[0051] This method is advantageous in that the structure can be loadedwith the therapeutic material in situ, i.e., at or near the point oftherapeutic use, typically before administration, preferablyimplantation, to a patient. This is particularly useful because thedevice can be stored and transported prior to incorporation of thetherapeutic material. This feature has several advantages. For example,the relevant consumer can select the therapeutic material to be usedfrom a wider range of therapeutic materials, e.g., a radioisotope with acertain half-life with certain particle emitting characteristics can beselected. Thus, the therapeutic material selected is not limited to onlythose supplied with the device but can instead be applied according tothe therapy required.

[0052] In another aspect of the invention, a sterilized structure isloaded with a therapeutic material. Preferably, the loading of thetherapeutic material includes contacting, more preferably immersing, theporous material in an aqueous solution comprising a salt of thetherapeutic material, wherein a water-insoluble salt of the therapeuticmaterial is formed within the porous material. Examples of therapeuticsalt solutions may be those previously mentioned above. The structure ispreferably packaged and can be shipped to the relevant consumer. Thestructure can now be administered to, preferably implanted into, apatient. Thus, in this embodiment, the structure is loaded with thetherapeutic material prior to reaching the point of use, which may bemore convenient depending upon the facilities available to the relevantconsumer.

EXAMPLES

[0053] The following nonlimiting examples will further illustrate theinvention. All parts, percentages, ratios, etc. are by weight unlessotherwise indicated.

Example 1

[0054] The following solutions were used in the procedure:

[0055] Solution A: 1-10% aqueous solution of BaCl₂

[0056] Solution B: 1-10% aqueous solution of Ba(NO₃)₂

[0057] Solution C: 1-10% aqueous solution of Na₂S³⁵O₄

[0058] A porcine fibrin stent made according to U.S. Pat. No. 5,510,077was treated by rehydration in Solution A by immersion for about 5 toabout 10 minutes. The stent was removed and excess solution was blottedwith absorbent paper. The stent was then dehydrated and sterilized bygamma radiation. Alternatively, Solution B can be used in place ofSolution A.

[0059] This treated stent was then rehydrated in an aqueous solution ofNa₂S³⁵O₄ radioisotope having a specific activity of about 10 μCi/ml toabout 500 μCi/ml. A white precipitate of BaS³⁵O₄ was observed within thepores of the stent surface. The stent can now be implanted into anartery for localized delivery of β-radiation or packaged for delivery tothe consumer.

Example 2

[0060] Fibrin stents made according to U.S. Pat. No. 5,510,077 weresoaked in a 20% by weight solution of BaCl₂ (preferably soaking forabout 10 to 30 minutes). The stents were then subjected to degassing byvacuum to remove air from the pores of the fibrin matrix, thus allowingthe BaCl₂ solution to fill the pores. The stents were dried overnight.The dried stents were placed into a solution of sodium heparinate(preferably soaking in a solution of 1000 U/ml to 20,000 U/ml for 10-20minutes—most preferably a solution of at least 10,000 U/ml) to allow theBaCl₂ in the fibrin matrix to react with the sodium heparinate to formbarium heparinate which was precipitated within the fibrin matrix.Scanning electron microscopy (SEM) showed that particulates of bariumheparinate on the order of 10 microns (i.e., micrometers) and smallerwere trapped within the fibrin matrix (FIG. 4). In vivo evaluation ofthe barium heparinate stents were carried out using a carotid crushmodel in pigs with standard fibrin stents as controls. After 24 hours,the stents were compared for flow and were then examined histologically.While flow did not differ in a statistically significant manner betweenthe control stent and the barium heparinate stent, the histologicalstudy showed substantially reduced clot formation on the lumenal surfaceof the barium heparinate stent (FIG. 5a) when compared with the controlstent (FIG. 5b).

Example 3

[0061] (A) Preparation of porous polyurethane coated stents. Wiktor typestents were placed over 3.0 mm diameter smooth glass rods and rolled byhand to assure a snug fit. The stent and rod assemblies were dipped in1-methy-2-pyrrolidinone (NMP) alone at room temperature, allowed todrain vertically for a few seconds, then rotated horizontally whiledusting with 400-500 mesh sodium bicarbonate until no furtherbicarbonate would adhere. After gently tapping the rod assemblies todislodge lightly adherent bicarbonate, the assemblies were dipped oncein a solution of 10 wt.% polyurethane in NMP. After draining verticallyfor a few seconds, the rod and stent assemblies were rotatedhorizontally while dusting with 400-500 mesh sodium bicarbonate until nofurther sodium bicarbonate adhered, then gently tapped to dislodgelightly adherent sodium bicarbonate. The assemblies were immersed inwater for about 5 minutes, then removed and the coating lightlycompacted by gently rolling the coated stent on the mandrel against awet paper towel. After immersing the stent assemblies in fresh water forat least 8 hours at room temperature the coated stents were removed fromtheir mandrels and immersed in fresh water for 4-8 hours at roomtemperature. The coated stents were subsequently dried in a forced airoven at 50° C. for about 8 hours and then trimmed of excess coatingbeyond the stent wires. After passing a visual inspection the porouspolyurethane stents were ready for subsequent fibrin impregnation.

[0062] (B) Preparation of the composite porous polyurethane-fibrinstent. The porous polyurethane stents prepared in (A) were suspended ina fibrinogen-thrombin solution. The ratio of fibrinogen to thrombin waspredetermined so that the clotting time of fibrinogen was between 5-10minutes. The fibrinogen-thrombin solution containing the porouspolyurethane stents was subjected to vacuum or ultrasonic degassing forabout 34 minutes. Using either method, the air in the pores of theporous polyurethane stents was driven out and the pores filled with thefibrinogen-thrombin mixture. The stents were then removed from thefibrinogen-thrombin solution and clotting of the fibrinogen in the poreswas further-carried out for another 10-20 minutes at room temperatureafter which the stents were incubated in sterile water overnight. Afterincubation, the stents were dehydrated at room temperature for at least2 hours. Alternatively, the porous polyurethane fibrin composite stentscan be further compressed in glass molds to densify the fibrin prior todehydration.

Example 4

[0063] In Vivo Assessment. A total of eight prototype stents fromexample 3 were implanted into 4 pigs for an in vivo assessment. Inparticular, a single stent was implanted into 2 coronary arteries ineach of the 4 pigs. The purpose of this pilot animal study was to asessthe deliverability of the composite (fibrin filled porous polyurethane)stents, the acute clinical performance, and to determine the 28-daybiocompatibility in terms of tissue response. The pigs were given 325 mg(ASA) per day throughout their course and heparin during the procedure.Stents were crimped onto commercially available PTCA catheters andimplanted using standard stent delivery techniques. All eight stentswere easily crimped tightly onto the balloons and successfully deliveredto the target site through 8 Fr or 9 Fr guide catheters. Implants weresuccessfully performed in all 3 major coronary arteries (RCA, LAD, LCX).Full expansion was acheived with 6-8 atmospheres pressure. Post-implantangiograms indicated good flow, with wide open lumens in all vessels.There were no acute events and the animals recovered without incident.Three of the animals reached their 28-day planned sacrifice date whilethe fourth pig was humanely sacrificed early at 21 days due to a chroniclung infection (non-cardiac related). All eight stented artery segmentswere pressure perfusion fixed and sectioned using standardhistopathology processing techniques, and subsequently examinedhistologically. Upon exam all eight stented artery segments were patent(i.e., open)—there were no total occlusions. In 3 of the pigs (6 stents)the lumens were widely patent with mostly thin to very thin amounts ofmaturing neointima deposited on top of the porous film. In addition,there was absence of or minor amounts of inflammation near the stentmaterial. In the 4th pig (2 stents) a thick neointima compromised thelumen with >60% stenosis and there was significant inflammatory responsein the tissue surrounding the stent material. In this particular animalit was noted that the arteries were small and there may have been anoverinjury component contributing to the gross vessel response. Thepores of the thin film in all these stents were typically filled withnative tissue as the exogenous fibrin appears to have been replaced witha cellular infiltrate typical of maturing neointima. There were alsosome areas of the porous film structure which were still containing theyet-to-be-absorbed exogenous fibrin.

[0064]FIGS. 6 and 7 are 28-day examples of elastic van Gieson (EVG) andhaematoxylin and eosin (H&E) stains respectively. FIG. 6 indicates theartery wall structure and expanded stent (wire holes plus porous film)with the thin film of neointima on top of the porous film. FIG. 7 is ahigher magnification photo of the cellular infiltrate into the porousfilm with neovascularization taking place at the porous film/neointimainterface. Original slide magnifications are 7x and 67x, respectively,for FIGS. 6 and 7. This data indicates that porous polyurethane stentshaving fibrin incorporated therein can be advantageously used forpreventing restenosis without any therapeutic material if so desired.

Example 5

[0065] (1) Porous polyurethane impregnated with barium heparinate. Aporous polymer stent prepared according to Example 3(A) was soaked in a20 wt. % solution of barium chloride (BaCl₂) for about 10-20 minutes.The stent was then removed and dehydrated for at least 2 hours(preferably overnight). After dehydration, the barium chlorideimpregnated porous polyurethane stent was placed in a sodium heparinsolution with a concentration of at least 10,000 U/ml. The stent wasimmersed in the heparin solution for about 10-15 minutes during whichtime the barium chloride reacted with sodium heparin to form bariumheparinate. The reaction took place within the pores containing bariumchloride and also at the surface of the stent where some of the bariumchloride leached out The barium heparinate, which is not soluble inwater, was trapped within the pores of the porous polyurethane stent.The stent was rinsed with sterile water to remove excess, unreactedbarium chloride and then dehydrated.

[0066] (2) Composite porous polyurethane-fibrin stent impregnated withbarium heparinate. The procedure used was similar to that in (1) aboveexcept that the stent described in Example 3(B) was used. After thereaction between barium chloride and sodium heparin was completed, thebarium chloride was replaced by barium heparinate that was trappedwithin the fibrin matrix.

Example 6

[0067] Preparation of porous polyurethane coated stent by using afreeze-immersion-precipitation method. Bare stents were cleaned with amixture of alcohol/water (50:50), and then Freon TE/TF (DuPont) and thendried. After cleaning, the stents were expanded to about 3.1 mm indiameter under clean room conditions. The stents were then individuallyplaced in glass mold cavities. The glass mold cavities have similarconfigurations as those described in U.S. Pat. No. 5,697,967. Apolyurethane solution (6.5 wt. % polyurethane in 1,4-dioxan) wasinjected into the mold cavities using a 3 ml sterile syringe and a 18 Gasterile needle. After injecting the polymer solution, the glass mold(containing the stents and polymer solution) was place in a refrigeratorat 3° C. for 2 hours. Since the freezing temperature of dioxan is at 12°C., the polymer solution frozen slowly thus creating a coarse structureof solvent/polymer. After 2 hours at 3° C., the mold was removed fromthe refrigerator and transferred to a freezer at −15° C. to −18° C. andkept at this temperature for an additional 1-1.5 hours. The mold wasthen immersed in an ice cold water bath at 3° C. for 34 days to allowthe solvent to leach out into the ice cold water. Porous, uniformpolyurethane stents were formed after the dioxan completely dissolvedinto the ice cold water. The mold (and the ice cold water bath) wereallowed to warm up slowly to room temperature. The stents were thenremoved from the glass mold cavities and continued to immerse in freshwater at room temperature for at least one day to remove traces ofsolvent. The stents were then air dried at room temperature under aclean room flow hood.

[0068] The above two-step freezing method was designed to create porouspolyurethane stents with desired pore structures and pore sizes. Byfreezing the polymer solution first at 3° C., large pores (in the orderof 50 micrometers or greater) were formed. This was reinforced withsmaller pores formed due to additional freezing and crystallization ofsolvent and polymer at −15° C. to −18° C. The two-step freezing methodcan create porous polyurethane stents with pore sizes range from 5 to120 micrometers (with average pore sizes of 50 to 60 micrometers). Poresizes at 50 to 60 micrometers are desirable for tissue ingrowth.

[0069] Fibrin and barium heparinate can also be incorporated in thistype of stent using the procedures described in Examples 3 and 5.

[0070] The complete disclosures of all patents, patent applications, andpublications referenced herein are incorporated herein by reference asif individually incorporated. Various modifications and alterations ofthis invention will become apparent to those skilled in the art withoutdeparting from the scope and spirit of this invention, and it should beunderstood that this invention is not to be unduly limited toillustrative embodiments set forth herein.

We claim:
 1. A medical device having at least one blood-contactingsurface comprising: a non-fibrin porous material having fibrinincorporated therein; and a water-insoluble therapeutic salt dispersedin at least a portion of the porous material.
 2. The medical device ofclaim 1 wherein the non-fibrin porous material comprises a film.
 3. Themedical device of claim 2 wherein the non-fibrin porous material isselected from the group consisting of silicone, polyurethane,polysulfone, cellulose, polyethylene, polypropylene, polyamide,polyester, polytetrafluoroethylene, and a combination of two or more ofthese materials.
 4. The medical device of claim 2 wherein thewater-insoluble salt comprises a salt of an antithrombotic material. 5.The medical device of claim 4 wherein the antithrombotic material isheparin.
 6. The medical device of claim 5 wherein the water-insolublesalt is selected from the group consisting of barium, calcium and silversalts of the antithrombotic material.
 7. An intravascular stentcomprising a generally cylindrical stent body and a single layer of apolymeric film; wherein the single layer of a polymeric film comprises anon-fibrin porous polymer with fibrin incorporated therein.
 8. The stentof claim 7 wherein the non-fibrin porous polymer film comprises amaterial selected from the group of silicone, polyurethane, polysulfone,cellulose, polyethylene, polypropylene, polyamide, polyester,polytetrafluoroethylene, and a combination of two or more of thesematerials.
 9. The stent of claim 8 wherein the non-fibrin porous polymerfilm comprises a polyurethane.
 10. The stent of claim 7 furthercomprising a water-insoluble therapeutic salt dispersed in at least aportion of the porous polymer film with fibrin incorporated therein. 11.The stent of claim 10 wherein the antithrombotic salt is a salt ofheparin.
 12. The stent of claim 11 wherein the water-insoluble heparinsalt is selected from the group consisting of silver, barium and calciumsalts.
 13. A method for coating a medical device with a porous polymer,the method comprising: placing the medical device in a mold; placing asolution of a polymer in the mold with the medical device; wherein thesolution of the polymer includes a solvent capable of phase separatingfrom the polymer at a temperature below the freezing point of thesolvent; cooling the solution of the polymer in the mold to atemperature below the freezing point of the solvent until a firstfraction of particulate material is formed by solidification and phaseseparation of the solvent from the polymer and is dispersed withinsolidified polymer; cooling the solution further and at a faster ratethan in the first cooling step to form a second fraction of particulatematerial dispersed within the solidified polymer, wherein the secondfraction of particulate material has a smaller particle size than thefirst fraction, and removing the particulate material from the polymerto form pores therein.
 14. The method of claim 13 wherein the medicaldevice is a stent.
 15. The method of claim 13 wherein the solutioncomprises ag polyurethane and dioxane.
 16. The method of claim 13wherein the porous polymer is in the form of a film.
 17. The method ofclaim 13 wherein the porous polymer is in the form of a coating.